The invention relates generally to systems and techniques for cooling magnetic resonance imaging systems and, more particularly, to a system for reducing the thermal energy transfer from the heated spots of an RF coil assembly to a patient bore.
Currently, Magnetic Resonance Imaging (MRI) systems have included a superconducting magnet that generates a temporally constant primary magnetic field. The superconducting magnet is used in conjunction with a magnetic gradient coil assembly, which is sequentially pulsed, to create a sequence of controlled gradients in the static magnetic field during an MRI data gathering sequence. The controlled gradients are effectuated throughout a patient imaging volume (patient bore), which is coupled to one or more radio frequency (RF) coils or antennae. The RF coils are located between the magnetic gradient coil assembly and the patient bore.
As a part of a typical MRI sequence, RF signals of suitable frequencies are transmitted into the patient bore. Nuclear magnetic resonance (nMR) responsive RF signals are received from the patient bore via the RF coils. Information encoded within the frequency and phase parameters of the received RF signals, by the use of an RF circuit, is processed to form visual images. These visual images correspond to the distribution of nMR nuclei within a cross-section or volume of the patient within the patient bore.
As is well known in the MRI industry, high power MRI systems consume large amounts of electrical power. In particular, the gradient and RF coils consume excessive amounts of power and, as a result, these coils generate significant heat. While a majority of the heat dissipation on the gradient coils is resistive in nature, the heat dissipation on the RF coil could be inductive due to the eddy currents induced by the time-varying gradient coil magnetic field. As a specific example, the high frequency switching during the running of gradient coil waveforms is known to cause eddy current formation at the end zones of the RF coils, thereby generating localized heat. As one would expect, excessive heat can cause system components to deteriorate or fail prematurely and hence adversely affects reliability. In addition, heat can be an annoyance to a patient during the imaging process. For this reason there are regulations that stipulate the maximum temperature of a patient support table, which effectively limits the amount of power that can be used in any MRI system. One way to minimize heat is to reduce coil currents but that solution reduces performance and can also adversely affect overall system efficiency.
Some MRI systems have been designed such that cooling air (i.e., chilled air) is passed through designated cooling air spaces between the RF coil and the gradient coils, thus dissipating coil heat. Unfortunately, designs of this type increase overall system volume, size, and costs. In addition, while air clearly reduces coil temperatures, in some cases the degree of cooling is insufficient to drive the coils at maximum coil currents and thus performance in these systems is minimized. That is, because of the limited heat flux entitlement associated with air cooling, air cooling is not completely adequate for cooling the RF coils and preventing the generated heat from transferring to the patient bore unless the air is chilled to a large degree. As a result, higher currents, cooling efficiency, and scanning times must be sacrificed in chilled air cooling systems in order to maintain the patient bore temperature within an acceptable, specified range thereby limiting performance of the MRI system.
One other solution has been to provide a hermetically sealed liquid cooling system with cooling conduits adjacent the gradient coils. According to systems having such a design, liquid coolant (e.g., water) is pumped through the system to cool the coils during field generation and data acquisition. Unfortunately, while liquid cooling systems have worked well for the purpose of cooling gradient coils, such systems have not been applied to cooling system components that reside inside the RF space such as the RF coils, the patient support table, etc. The primary reason for not providing a liquid cooled configuration that extends into the RF space is that coolant hydrogen atoms, like human tissue, include a large number of protons that, when inside the RF shield, tend to generate nMR signals. These spurious signals, like the signals generated by the excited human tissue, are received by the detector coils and distort the resulting data and associated images. Thus, liquid cooling systems have been limited to areas outside the RF shield to avoid spurious signal excitation and air cooling systems have been employed for cooling the RF coils.
It would therefore be desirable to have a system and method that minimizes thermal energy transfer from the RF coil assembly to the patient bore without degradation of nMR signals.